The present invention relates to a multi-channel semiconductor radiation detector for detecting radiation (gamma rays) emitted from radioactive isotopes (radioisotopes, RIS) injected into a subject such as a patient, and a nuclear medicine diagnostic apparatus for generating the RI distribution in the body on the basis of the detection result.
A nuclear medicine diagnostic apparatus can detect radiation such as gamma rays emitted from a subject into which RIs are injected, and obtain the RI distribution in the body on the basis of the detection result. With this RI distribution, fundamental data of functional diagnoses, e.g., identification of a morbid portion in the body and measurements of the blood flow rate and the fatty acid metabolic amount can be acquired.
As this nuclear medicine diagnostic apparatus, a SPECT apparatus using single photon emission computer tomography (SPECT) and a PET apparatus using coincidence detection type positron emission computer tomography (PET) are known. The latter PET apparatus includes a plurality of detectors and performs imaging by simultaneously detecting gamma rays emitted at an angle of 180.degree. when positrons combine with electrons and disappear.
Recently, a SPECT apparatus which includes a plurality of detectors to perform both the SPECT and the coincidence detection type PET is becoming popular. A "nuclear medicine diagnostic apparatus" is a general term for these apparatuses.
As a detector of the conventional SPECT apparatuses, an Anger type detector in which a plurality of photomultiplier tubes are densely arranged on a scintillator is used most frequently. However, this Anger type detector is large and has relatively low energy resolution and detection characteristics.
A prevalent method used in the coincidence detection type PET apparatus is to perform imaging by simultaneously detecting the timings at which gamma rays are emitted at an angle of 180.degree. by using the combination of a bismuth germanium oxide (BGO) detector and a photomultiplier or a photodiode.
Also, a nuclear medicine diagnostic apparatus which uses an Anger type detector having a plurality of detectors and can realize the coincidence detection type PET by mode switching is currently put to use in many cases.
In any nuclear medicine diagnostic apparatus, however, a detector for detecting gamma rays is a scintillator type detector. So, it is necessary to once convert incident gamma rays into weak light by the scintillator and convert this weak light into an electrical signal by a photomultiplier tube or the like. This increases the size of a nuclear medicine diagnostic apparatus and limits its performance.
A semiconductor radiation detector, therefore, is attracting attention. Since a semiconductor radiation detector converts gamma rays directly into an electrical signal, the efficiency of conversion to an electrical signal is high. Additionally, semiconductor detecting elements can individually detect gamma rays. Hence, a semiconductor radiation detector is expected to improve the energy resolution and detection characteristics.
A cadmium telluride (CdTe) semiconductor cannot form a single-crystal structure such as formed by sodium iodide (NaI) used in the present Anger type detector. In one prior art, therefore, a two-dimensional semiconductor radiation detector is constructed by densely mounting small detector modules (incorporating a two-dimensional semiconductor array cell and a preamplifier and read circuit formed below this two-dimensional array cell so as not to extend from the cell). Unfortunately, interconnections between these detector modules and dead spaces inside the detector modules are nonuniform, and a unique artifact occurs. This makes reconstruction of RI images difficult.
Additionally, a modular construction like this can perform only signal processing each detector module is capable of. Therefore, signal processing is presently very difficult when coincidence detection is observed over a plurality of semiconductor detecting elements (i.e., a plurality of detector modules) in a nuclear medicine diagnostic apparatus, such as the coincidence detection type PET apparatus, which processes a high energy of 511 kev.
Also, it is difficult to form fine pixels such as in a conventional digital gamma camera by using semiconductor detecting elements used in a semiconductor radiation detector due to restrictions on the cost and packaging method. This limits the formation of fine structures.
The Anger type detector is generally designed as a plane detector. As shown in FIG. 1, an Anger type plane detector has a plane parallel collimator 10 and a plane scintillator 11. An RI injected into a subject P emits gamma rays from inside the body of the subject P, and the scintillator 11 converts the gamma rays into light via the parallel collimator 10. This light is converted into an electrical signal by, e.g., a photomultiplier tube (not shown) and subjected to signal processing. The subject P has a curved surface. Therefore, although the center of the Anger type plane detector shown in FIG. 1 is close to the subject P, this plane detector separates from the subject P toward the periphery of the subject P. Consequently, the positional resolution pertaining to detection of the incident position of gamma rays generally deteriorates.
Instead of this Anger type plane detector, a detector having curved surfaces which allow the whole detector to be located close to a patient can be used. As shown in FIG. 2, this Anger type gamma camera (manufactured by BICRON Corp.) includes two concave curved detectors 20a and 20b opposing each other with a subject P between them. These curved detectors 20a and 20b are separated by a rotation radius R from a rotation center O and moved around the subject P along a rotating direction D1 by a driving mechanism (not shown).
The curved detectors 20a and 20b have the same arrangement; the curved detector 20a includes a scintillator 21a, a light guide 22a, and photomultiplier tubes 23a to 23n, and the curved detector 20b includes a scintillator 21b, a light guide 22b, and photomultiplier tubes 24a to 24n. Each of the scintillators 21a and 21b is a concave cylindrical member and has a fixed thickness t1 in a direction to the rotation center O. The photomultiplier tubes 23a to 23n and 24a to 24n opposing the rotation center O detect the incident position of gamma rays.
In the curved detector 20a or 20b with the above construction, the accurate incident positions of gamma rays entering the center and vicinity of the scintillator 21a or 21b can be calculated by using the photomultiplier tubes because there is no big difference between the incident positions. However, if parallel collimators (not shown) are set on the subject sides of the scintillators 21a and 21b to detect parallel gamma rays in the SPECT, the distance the gamma rays travel through the scintillator 21a or 21b increases from the center to the peripheries of the curved detector 20a or 20b. This is so because the scintillators 21a and 21b have the thickness t1 in the form of a cylinder in the direction to the rotation center O.
Accordingly, the influence of DOI (Depth Of Interaction) increases, and this produces a positional resolution error .DELTA. in accordance with the position where the scintillator 21a or 21b absorbs gamma rays. Consequently, the accuracy of calculated positional resolution deteriorates. Although the peripheries of the curved detectors 20a and 20b are located close to the subject P, the positional resolution does not increase. This prevents the SPECT from fully utilizing the merit that the peripheries of the detector are closely located to a subject. Additionally, the sensitivity changes in accordance with the incident position of gamma rays (the sensitivity rises toward the peripheries). Hence, it is difficult to increase the sensitivity in the center, which is originally preferably high.
In the coincidence detection type PET, the position of a positron PO (511 keV) along a line extending perpendicularly to the curved surfaces of the scintillators 21a and 21b from the rotation center O is ideally calculable. However, when parallel gamma rays are detected by using parallel collimators as described above, the distance the gamma rays from the positron PO travel through the scintillator 21a or 21b increases toward the peripheries to increase the influence of DOI, because the scintillators 21a and 21b have the fixed thickness t1 in the form of a cylinder in the direction to the rotation center O. Since this deteriorates the positional resolution, the quality of a reconstructed RI image deteriorates.
Especially when the thickness t1 of the scintillators 21a and 21b is large, an incident position where the influence of DOI is conspicuous exists depending on the incident angle of gamma rays from the positron PO. For this reason, when a gamma camera in which two detectors oppose each other is used, satisfactory image quality can be obtained only when these two detectors are separated a predetermined distance or more, because a coincidence detection solid angle .theta.1 becomes small. Accordingly, although a curved detector originally has the merit that its solid angle is larger than that of a plane detector, the coincidence detection solid angle .theta.1 for gamma ray detection in the PET cannot be increased more than expected of a curved detector. As a consequence, the sensitivity cannot unlimitedly increase.
Additionally, for the reasons described earlier, the positional resolution largely deteriorates in the peripheries of a detector even in the fan beam SPECT in which the focal length is short.
Furthermore, when the simultaneous use of the SPECT and the coincidence detection type PET is desirable, it is difficult to perform these two methods at the same time and obtain a greater merit than when a plane detector is used.
Also, as described above, the scintillator has a fixed thickness in the form of a cylinder in the direction to the rotation center 0. Therefore, when parallel gamma rays are detected by using parallel collimators, the sensitivity changes in accordance with the incident position of the gamma rays (the sensitivity rises toward the peripheries). Consequently, it is difficult to increase the sensitivity in the center of the detector, which is originally preferably high.
Moreover, to acquire projection data concerning cardiac muscles over an angle of 180.degree. by using the Anger type plane detector, there is a method by which two plane detectors are separated by an angle of 90.degree. and the whole detector assembly is rotated 90.degree.. Since, however, the incident surface of each detector is a plane surface, these incident surfaces cannot approach the heart of a subject more than expected.